This invention relates to an improved system for radiographic imaging and material analysis and more specifically for nuclear medicine and mammography imaging.
Two general imaging problems in radiology involve the determination of a radiation source distribution and/or the effect of a filter, in this case a patient, on the radiation source distribution. Consider the field of nuclear medicine where the radiation source or other radionuclide distribution emits photons or positrons, Image data acquisition in nuclear medicine presents several challenges in addition to constraints imposed by finite acquisition times and patient exposure restrictions. Most photon energies that are of interest in nuclear medicine are higher than the typical photon energies employed in diagnostic x-ray radiography. In particular, Positron Emission Tomography (PET) involves the detection of pairs of very high energy photons due to annihilation events. Unfortunately, the photon radiation source, such as a radionuclide, used in nuclear medicine is not directional and the source distribution within the body is not precisely known.
Photons that escape the body may be scattered, altering their energies and/or direction vectors. It is desirable for many applications to discriminate against scatter radiation reaching the detector based on energy and/or direction. It may also be desirable to only detect radiation with a specific direction vector, since many detection systems possess poor directional discrimination capability and have finite response times within which to detect events, thereby limiting detection rates. Thus detection systems used in nuclear medicine such as Gamma cameras or PET scanners often employ conventional, such as attenuating or rigid geometry, focused or unfocused collimators, often referred to as grids or grid collimators, to help define the direction vectors of a detected photons. The direction vectors and energies of non-scattered photons are well-defined. Unfortunately, the emission of photons from the source distribution is non-directional and the radiation source distribution itself is typically not well-defined. A Compton-scattered photon suffers an energy loss and change in direction vector whereas a coherent or Rayleigh scattered photon only has its direction vector altered. In general x-ray radiography the source is a x-ray tube, although a radionuclide maybe substituted, used in a point, slit, slot, or area imaging configuration. The energy distribution and direction vector of the radiation from a x-ray tube are approximately known. These parameters are typically well-defined for a collimated radionuclide source used in an application such as point-scan Compton scatter imaging and material analysis. A number of detection. formats are in use depending on the application. A planar detector geometry is typically utilized for applications such as mammography, angiography, and chest radiography which typically employ detectors such as x-ray film-screen devices, or storage phosphor screens, or image intensifiers coupled to cameras. Slit- and slot-scan formats are also available, usually incorporating improvements to the detectors and, in some instances, the radiation source. Additional image acquisition formats include ring-shaped detectors or flat detectors for fan-beam or cone-beam tomography, respectively. Common detector geometries used in nuclear medicine typically include one or more planar detectors, which are basically standard Gamma cameras, with attached conventional collimators or ring detectors, used in Positron Emission Tomography. Imaging systems based on standard Gamma camera and related detector designs are frequently used for a number of nuclear medicine studies such as heart; brain, thyroid, gastro-intestinal, whole body, and breast imaging, including scintimammography. A basic Gamma camera design employs a large, planar array of scintillation crystals or a single, large, planar scintillation crystal optically coupled to an array of photomultiplier tubes (PMTs). A conventional focused or unfocused collimator is typically mounted to the face of the Gamma camera. This inflexible imaging system is then positioned such that the region of interest containing the source distribution is within the field of view. It provides a limited degree of spatial resolution and energy resolution while removing some fraction of unscattered and scattered radiation that would otherwise degrade image quality. Unfortunately a substantial fraction of useful unscattered radiation is also attenuated. Another infrequently used design replaces the conventional collimator with a coded aperture such as a uniformly redundant array aperture which is also based on photon attenuation and is typically rigid. Commercial systems may use one, two, or three Gamma camera detector units. One commercial system eliminates the use of scintillator crystals and PMTs with a rigid, planar, 2-D CdZnTe semiconductor detector manufactured by abutting four 2-D CdZnTe arrays of moderate size. Techniques for abutting 2-D silicon arrays are well-known in the art. Drawbacks to employing large- or medium-sized 2-D CdZeTe arrays capable of high detection efficiency include the difficulty of growing thick CdZnTe crystals with acceptable levels of defects and creating a low noise, 2-D array readout structure on top of a large- or medium-size CdZnTe crystal. Grid collimators are still desirable for many applications since the direction vectors of detected photons are otherwise poorly defined. A design which replaces a conventional collimator with a relatively thin, planar semiconductor, often Ge, array of moderate size, which serves as a Compton scatterer is referred to as a Compton electronic Gamma camera. This system is still being refined. The detector module array described below can be used in place of a standard Gamma camera in a Compton Gamma camera system.
Nuclear medicine imaging applications are complicated by the fact that the spatial distribution of the source within a region of the patient is poorly defined. One way to simplify this problem is to use emitted photons of known energies. For example, a source that has one or more emission energies of a narrow energy bandwidth may be utilized. The problem now is the reconstruction of the source distribution rather than the calibration of the source distribution. The measured source distribution, i.e., the apparent source distribution, represents the filtered true source distribution, assuming self attenuation is small. In certain nuclear medicine applications estimates of the true source distribution are obtained by calibrating the contribution of the filter, which may be the patient, to the apparent source distribution. Photon transmission measurements are made in order to estimate the effect of tissue scattering and absorption or attenuation on radiation source measurements by using a reference source that is external to the patient. Unfortunately, measuring photon transmission through the body does not duplicate the actual imaging chain acquisition format used in nuclear medicine where photon are transmitted out of the body. Photons in the two instances do not traverse comparable paths.
In accordance with the present invention, a radiation detection apparatus is provided for radiographic imaging and material composition analysis in which the apparatus can dynamically configure its array geometry and radiation detector parameters for a specific imaging task or it can use an existing radiation detection geometry and settings. This invention is particularly suited for x-ray and gamma ray imaging in nuclear medicine, including scintimammography, and x-ray radiography, specifically, x-ray mammography. There are several advantages inherent to this invention. Superior detectors in cost-effective formats can be utilized and detectors with different properties, including materials, resolution, response time and noise characteristics, can be used within an array. One or more radiation detectors are incorporated into a detector module and one or more modules make up a detector module array. The detector modules transmit detected photon image data and relevant module parameters to a computer system which utilizes this information to electronically-control the modules and in some cases attached collimators. This system is implemented using detector sub-arrays, comprised of one or more detector modules, and detector arrays in order to enhance image quality or analysis capability. Conventional attenuating or rigid geometry collimators, including ones characterized by coded apertures, and unconventional, including x-ray optic, configurable (adaptive), and Compton scatter module, collimators can be employed to improve the energy and/or spatial resolution for the photon radiation detection system. In a similar manner additional types of radiation optic collimators such as neutron optic collimators or electron optic collimators capable of focusing electric or magnetic fields, can be used with neutrons or charged particles, respectively.
In a preferred embodiment semiconductor detectors with appropriate geometries, such as edge-on detectors; thick, linear array detectors; or small, thick, 2-D array detectors, are incorporated into detector modules which are mounted within a frame and configured as an array of detector modules. Detector modules contain one or more detectors, possibly with different properties. A detector array contains one or more modules or types of modules. For nuclear medicine imaging applications detector sub-arrays, comprised of one or more modules, or the entire detector array can be positioned and oriented with respect to the radiation source by an operator or by direct computer control. Collimators and shielding can be attached to or integrated into the module, including interfacing with module electronics if appropriate. Modules communicate with the computer system which monitor and control module and collimator parameters and collect and process radiation data recorded by the detectors. Modules may communicate directly or through a shared network with the computer system. Computer-controlled services include sending electronic instructions to the module mounting hardware, the module, and the collimators, if appropriate. Electronic instructions can initiate actions such as detector array motion, adjustment of the relative position or orientation of one module with respect to other modules, manipulation of a collimator, and the modification of module operating parameters, such as detector signal amplification, filtering, resolution, temperature, operating voltage or sampling rate. Since positioning machinery can be incorporated into the module, actuators can be employed to adjust the position and orientation of the detector. The actuators can also manipulate the positions and orientations of appropriate collimators. A novel collimator design utilizes actuators to alter the configuration of a collimator. The computer-based monitor and control capabilities can be used to track and adjust the locations of modules while they are in motion. Positions, orientations, and motion of all detectors and relevant collimators are recorded and updated as needed throughout the image acquisition process.
A typical nuclear medicine imaging session begins with an operator selecting from a computer display menu a specific detection system with pre-defined array geometry, collimator, and module settings appropriate for the desired imaging task. The detector array configuration can already exist or it can be set up by the computer system. Once a baseline detection system is established, an operator can then adjust and fine tune the detector array position and settings or leave the detector array adjustments and tuning under computer control. While under computer control electronic instructions can be issued dynamically in response to detector module parameter values and detected radiation data that is transferred to the computer system for processing, display, and storage during image acquisition or adaptive imaging. Electronic commands can be used to control the array geometry and motion, detector module parameters, and some types of collimators. Thus an information feedback loop can be implemented as a means of tuning detection system parameters. For some imaging or analysis applications it will be sufficient to configure the detector array based on either a standard geometry, such as line, plane, open box, wedge, ring, cylinder, ellipse, ellipsoid, sphere, or a contoured geometry, in order to compensate for the radionuclide distribution within the subject and/or the shape of the subject at the region of interest. For example, configurations may be based on the breast size of a woman or on the head size, waist size, or chest size of an adult, child, or infant. A versatile design allows at least a subset of these detector array geometries to be generated xe2x80x9con the flyxe2x80x9d. A less-versatile design still utilizes modules, but the modules are fixed within a specific detector array geometry or they are constrained to move to specific positions, for example, along a track, within a specific detector array geometry. Less-versatile designs reduce the mechanical complexity of the detection system and may be sufficient for specific imaging tasks. An optional capability is to allow the entire array to undergo discrete or uniform motion. The simplest example of, this capability would be to scan a radiation source with a detector array comprised of a single detector module.
In another embodiment, semiconductor detectors are replaced by other types of suitable detectors, such as scintillation detectors, gas detectors, liquid detectors, or superconducting detectors.
In another embodiment reference sources are introduced into the subject and then imaged. The size, shape, intensity, and emission spectrum of the reference sources are known. This allows measurements to be made of photon attenuation due to material in the photon path prior to reaching the detector. This information can be used to estimate the true source distribution from measurements of the apparent source distribution made during image acquisition in a nuclear medicine test. The reference source can also be used to focus the detector array in order to tune the imaging chain.
In another embodiment detector modules and collimators are incorporated into x-ray radiography slit scan imaging systems. X-ray optic collimators can be used to increase the intensity and modify the spectrum of the x-ray radiation that is recorded by the detector module. A single x-ray source is combined with a x-ray optic collimator and a x-ray detector module and used for a x-ray mammography slit scan system. Another improvement involves aggressively compressing sections of the breast and acquiring separate images of the highly-compressed sections rather than acquiring a single image of the entire, mildly-compressed breast.
The system of the present invention may utilize devices detailed in prior inventions for slit-scan or slot-scan radiographic x-ray imaging in which photons are detected directly using edge-on array detectors; small, 2-D semiconductor array detectors; or semiconductor array detectors coupled to scintillators. This new device can also use thick, linear semiconductor array detectors and thick, small, 2-D semiconductor array detectors in addition to other types of detectors. Manufacturing costs for these detectors are much less than those associated with large-area or moderate-area, thick, planar, 2-D semiconductor array detectors made from materials such as, but not limited to, CdZnTe, CdTe, GaAs, Ge, Si, SiC, or HgI2. The detector format is also compatible with detectors such as thin, linear semiconductor arrays or thin, small 2-D semiconductor arrays optically-coupled to scintillators. For example, thin, linear semiconductor arrays of avalanche photodiodes optically-coupled to scintillators can be used as radiation detectors. Another example of a radiation detector is an edge-on scintillator detector that consists of a scintillator array optically-coupled to a semiconductor detector which provides the readout signal. This approach can be extended to include scintillators coupled to integrated photoemissive cathodes or small PMTs; small, gas microcapillary detector assemblies or small superconducting array detectors. Consider a scenario in which radiation is incident upon a planar edge-on detector. The detector thickness (height) now defines the maximum detector entrance aperture while the length or width of the detector area now defines the maximum attenuation distance for edge-on radiation detector designs including semiconductor drift chamber, single-sided strip, and double sided strip detectors, including micro-strip detector versions. Strip widths can be tapered or curved, in the case of drift chamber detectors, if focusing is desired. In the case of double-sided parallel strip detectors in which opposing strips are parallel, both electrons and holes can be collected to provide 2-D position information across the aperture. If strips on one side run perpendicular to those on the other side, then depth-of-interaction information can be obtained. If strips are segmented in either a single-sided or double-sided parallel strip detector then depth-of-interaction information can be obtained and readout rates can be improved.